Early in this century, the Austrian mathematician J. Radon demonstrated that a two-dimensional slice of a three-dimensional object may be reproduced from an infinite set of all of its projections. Computed tomography ("CT") X-ray systems generate an infinite set of X-ray beam projections through an object to be examined. The resultant detected X-ray data are computer processed to reconstruct a tomographic image-slice of the object.
More specifically, CT systems subject the object under examination to one or more pencil-like X-ray beams from all possible directions. X-ray beams passing through the object are attenuated by various amounts, depending upon the nature of the object traversed (e.g., bone, tissue, metal). One or more X-ray detectors, disposed on the far side of the object, receive these beams and provide analog output signals proportional to the strength of the incoming X-rays. Each detector output is then digitized and computer processed to help produce an image of a slice of the object.
For an ideally reconstructed image, a CT system would have an infinitely large number of extremely tiny closely-spaced detectors (or fewer detectors repositionable to an infinite number of very closely spaced locations). Because of limitations imposed by economics and hardware, practical CT systems process X-ray beam data from hundreds or thousands of detectors (or detector positions). However, the cost of analog-to-digital ("A/D") conversion of the output of hundreds or thousands of detectors can be prohibitive. The present invention discloses an apparatus and method for reducing the complexity of A/D converters required by CT systems, as well as the number of digitized information channels to be computer processed, without substantial degradation of the image produced.
At this juncture it is helpful to briefly overview the various types of CT X-ray scanner systems that have been developed, to appreciate better the problems associated with obtaining sufficient data to provide a meaningful output image.
A "first generation" CT scanner configuration is schematically depicted in FIG. 1A, and includes a single X-ray source (S), and a single detector (D), both of which move in a translate-rotate fashion (indicated by arrows T, R) relative to the object under examination (O). The translate-rotate movement causes the X-ray beam B to pass through the object O from all directions, with the moving detector D providing an analog output signal at many detector positions responsive to the various beam paths. A signal processing mechanism (SP) digitizes the output of D and ultimately reconstructs an image of a slice of the object O.
For ease of illustration, FIG. 1A (and indeed FIGS. 1B, 1C, 1D) depicts detector D and source S rotating with an equal radius about the object O, but in practice these radii need not be equal. Further, as is also the case with FIGS. 1B, 1C and 1D, relative motion between the object and the source and detector is all that is required, and in some applications the detector D and source S are stationary and the object O is rotated. For example, some industrial CT scanners that image large objects (an airplane wing, for example) require a source S which is too large to move, and therefore instead move the object relative to a stationary source and detector.
FIG. 1B depicts a second generation scanner wherein a source S emits an X-ray fan beam B, portions of which pass through the object O and are detected by a single detector D, where both source and detector move in a translate-rotate fashion. Because a fan beam of X-rays is used, the system of FIG. 1B can complete a scan more quickly than the system of FIG. 1A. For example, if the fan beam B comprises N rays, object O may be scanned N times faster than the first generation system of FIG. 1A wherein there is but one ray. For ease of illustration, FIG. 1B depicts a small number of rays, but fan beam B actually includes a continuum of rays. Again, it is understood that detector D and source S need not be equidistant from the object O, and that relative motion between the object and the source and detector, is all that is required. A signal processing mechanism (SP) digitizes the detector output and provides an image of a slice of the object O.
FIG. 1C depicts a third generation CT scanner system having a point X-ray source S and multiple detectors D.sub.i arranged in an arc, where point source S and the array of detectors D.sub.i rotate relative to the object O, but there is no translational movement. Conceptually if the single detector of FIG. 1A could be relatively rotated in an arc to various detector positions corresponding to D.sub.i, it could serve as the plurality of detectors shown in FIG. 1C. Again, the radii of relative motion are drawn equal for ease of illustration only. A signal processing mechanism (SP) digitizes the detector outputs and provides an image of a slice of the object O. U.S. Pat. No. 4,630,202 to Mori (Dec. 16, 1986) describes a conventional third generation CT scanner.
FIG. 1D depicts a fourth generation CT scanner system as including a source S that emits an X-ray fan beam B and an array of detectors D.sub.i, wherein the source S rotates relative to the object O, the detector array is stationary, and no translational movement of source or detectors occurs. Again it is understood that the detector array and X-ray source are drawn as being equidistant from the object 0 merely for ease of illustration. A signal processing mechanism (SP) digitizes the detector outputs and provides an image of a slice of the object O.
FIG. 1E depicts in greater detail a fourth generation scanning electron beam CT system such as described generally in U.S. Pat. No. 4,352,021 to Boyd, et al., (Sep. 28, 1982). In the CT system 10 of FIG. 1E, there is a housing chamber 12 wherein an electron beam 14 is generated and caused by a beam optics assembly 16 to scan an arc-like typically tungsten target 18 located within chamber 12's front lower portion 20. Upon being struck by the electron beam, the target emits a moving fan-like beam of X-rays 22. At least some of these X-ray beams pass through a region of object 24 and then register upon a region of a detector array 26 located generally diametrically opposite that portion of target 18 struck by the electron beam 14. The detector array includes a relatively large number of detectors D (perhaps 2,000-5,000) that sequentially receive at least a portion of the moving beam of X-rays. Each detector's output is an analog signal typically representing a few nA of current, which signal is passed through a current-to-voltage converter (not shown) to yield an analog signal of perhaps several hundred mV. Preferably each detector output is lowpass filtered (filters not shown) to remove frequency components higher than about 25 KHz. The analog detector outputs are then digitalized using analog to digital ("A/D") converters.
To reduce the number of A/D converters required (and thus the number of digitized signal channels requiring computer processing downstream), blocks of analog detector outputs are typically multiplexed. The multiplexed outputs are then digitized using A/D converters, with fewer A/D converters required because of the multiplexing. These multiplexing and converting steps are carried out by the Digital Acquisition System (DAS) depicted generally in FIG. 1E by box 30. The digitalized signal channels from the A/D converters are then coupled to a computer system 32 where they are extensively processed to produce an image of a slice of the X-rayed subject on a high resolution video monitor 34. Typically computer system 32 also controls reading of the various detectors, electron beam scanning, and repositioning of the X-rayed subject 24. With reference to FIGS. 1A-1D, the signal processing mechanism (SP) depicted therein may be thought of as including elements 30, 32 and 34 as depicted in FIG. 1E.
FIG. 1E helps demonstrate that in reconstructing a tomographic image from X-ray detector data, it is unimportant how the X-ray beams were caused to traverse the object O. For example, while FIG. 1E depicts a scanning electron beam system, one could just as easily mount an X-ray generator on a mechanical gantry and rotate the gantry (which could include the detector array) about the subject 24. Such a scanner system is in fact depicted in U.S. Pat. No. 4,630,202 to Mori (Dec. 16, 1986), which describes a third generation system. Similarly, in the other configurations described above, the moving source of X-ray beams or fan beams could be produced by rotating an X-ray source of such beams, or by scanning a stationary X-ray emitting target with a moving electron beam (as is done in FIG. 1E).
FIG. 2 is a detailed depiction of the contents of the DAS 30 in FIG. 1E, and depicts a prior art technique for reducing the number of A/D converters and resultant digitalized information channels that require computer processing. FIG. 2 also depicts several relationships required to reconstruct an image from CT data.
In the center of FIG. 2 is shown the isocenter 21 for a scanner system 10. By definition, the isocenter is the center of a fictitious image reconstruction circle 23, within whose circumference must lie the object O to be scanned and suitably reconstructed. Further, isocenter 21 is also the center of the array 25 of the detectors D.sub.i, and the center of the X-ray beam source or scan path 27 (which coincides with target 18 in the embodiment of FIG. 1E). For acceptable reconstruction, the region within the circle 23 must be exposed to X-ray beams from all possible directions.
The depiction of FIG. 2, wherein the vortex of the beam fans shown is a source point is termed a "source fan". In a source fan depiction, it is understood that at any given time, an X-ray fan beam simultaneously strikes many detectors. Although for ease of illustration only two X-ray fan beams are shown, each having three beams, a fan beam such as B.sub.t is actually a continuum comprising a large number of X-ray beams.
The relationship between the isocenter, the image reconstruction circle, the location of the detector(s) and X-ray source or X-ray beam source exists for any of the systems shown in FIGS. 1A-1E. While FIG. 2 depicts the X-ray source or X-ray beam source as being radially more distant from the isocenter than the detectors, the converse could be true, or in a suitable arrangement equal distances could be employed. Any of the CT systems shown in FIGS. 1A-1E can suitably reconstruct any image 0 lying within the reconstruction circle 23. Note in FIG. 2 that the lower ends of the detector array 25 extend at least to (and preferably slightly beneath) the lower edge of the reconstruction circle 23, while the upper ends of the scan path 27 extend to (and preferably slightly above) the upper edge of the reconstruction circle 23. This configuration ensures that any point within circle 23 is scanned from every possible direction (including horizontally), and that all X-ray beams fall at least partially on a portion of the detector array 26.
Detector array 26 in FIG. 1E preferably may include as many as 4,800 detectors (or detector positions) spaced apart a distance d about 1 mm or 0.1.degree. (see FIG. 3). However for ease of illustration only nine such detectors (D.sub.1, D.sub.2, . . . D.sub.9) are depicted in FIG. 2, disposed along arc 25. Because the X-ray fan beam 22 is depicted by arrow 28 as moving clockwise, it is understood that detector D.sub.1 is scanned first by the beam 22, then detector D.sub.2, D.sub.3 and so forth. For example, at an arbitrary time t, a fan beam B.sub.t is emitted at a position S.sub.t along the tungsten target 18 (see FIG. 2). A short time later, the electron beam 14 (see FIG. 2) scans position S.sub.t+1 along the target 18 and a fan beam B.sub.t+1 results. The result is a scanning fan beam B that sweeps clockwise along arc 27 in FIG. 2. Depending upon mode of operation, for the system of FIG. 1E the electron beam 18 preferably scans about 210.degree. along the target 18 in perhaps 50 mS to 100 mS.
For ease of illustration, FIG. 2 does not depict the current to voltage conversion ("I/V") for each detector output, or any filtering to limit higher frequency detector output components. At any given time, only detectors exposed to the moving X-ray fan beam 22 need be activated or read. For a system such as that shown in FIG. 1E having about 2,000 detectors, a given detector is activated and its output sampled every 10 .mu.s to every 50 .mu.s or so. This selective detector activation, preferably controlled by computer system 32, promotes efficient signal processing because time is not wasted sampling detectors having no meaningful output to contribute to the reconstructed image.
Because they are not simultaneously read or A/D converted, detectors D.sub.1, D.sub.2, and D.sub.3 may be multiplexed together using multiplexer 40, whose single analog output is digitized with an A/D converter 42. Similarly, outputs from detectors D.sub.4, D.sub.5, D.sub.6 and D.sub.7, D.sub.8, D.sub.9 may be multiplexed using multiplexers 44 and 48, respectively, whose single outputs are then digitized by A/D converters 46 and 50, respectively. It is of course not necessary that the multiplexed detectors be adjacent to each other, only that they not be simultaneously converted.
Thus, in the example depicted in FIG. 2, by multiplexing the output from three detectors, the number of required A/D converters is reduced to one-third. In a CT system with about 4,800 detectors, the use of multiplexers can reduce the number of required A/D converters from about 4,800 to about 1,700. Unfortunately providing a CT system with 1,700 A/D converters is still rather expensive, and still represents approximately 1,700 channels of digitized signal data to be computer processed downstream.
Different multiplexing schemes can of course reduce the number of A/D converters (and thus the number of digitized information channels requiring processing) by a factor greater than three. However aside from such multiplexing techniques, it is not known in the art how to further reduce the number of A/D converters without substantial degradation of the resultant CT system video image. For example, conventional wisdom has long considered it unfeasible to directly combine the outputs of two or more detectors without loss of image resolution due to degradation of the detector modulation transfer function ("MTF").
To recapitulate, what is needed is an apparatus and method for combining detector output signals in a CT system that not only reduces the number of information channels to be processed (using multiplexers) but also reduces the complexity of the A/D conversion circuitry and the associated delay and combining circuitry. Additionally, it is important that such apparatus and method should not substantially degrade the resultant video image, should be relatively inexpensive to implement, and preferably should operate in real time in a conventional CT system. The present invention discloses such an apparatus and method.